Radio-frequency coil signal chain for a low-field mri system

ABSTRACT

Methods and apparatus for reducing noise in RF signal chain circuitry for a low-field magnetic resonance imaging system are provided. A switching circuit in the RF signal chain circuitry may include at least one field effect transistor (FET) configured to operate as an RF switch at an operating frequency of less than 10 MHz. A decoupling circuit may include tuning circuitry coupled across inputs of an amplifier and active feedback circuitry coupled between an output of the amplifier and an input of the amplifier, wherein the active feedback circuitry includes a feedback capacitor configured to reduce a quality factor of an RF coil coupled to the amplifier.

RELATED APPLICATIONS

This application claim priority under 35 U.S.C. § 119(e) to U.S.Provisional Application No. 62/674,458, filed May 21, 2018, and titled,“Radio-frequency Coil Signal Chain for a Low-field MRI System,” and U.S.Provisional Application No. 62/692,454, filed Jun. 29, 2018, and titled,“Radio-frequency Coil Signal Chain for a Low-field MRI System,” theentire contents of each of which is incorporated by reference herein.

BACKGROUND

Magnetic resonance imaging (MRI) provides an important imaging modalityfor numerous applications and is widely utilized in clinical andresearch settings to produce images of the inside of the human body. Asa generality, MRI is based on detecting magnetic resonance (MR) signals,which are electromagnetic waves emitted by atoms in response to statechanges resulting from applied electromagnetic fields. For example,nuclear magnetic resonance (NMR) techniques involve detecting MR signalsemitted from the nuclei of excited atoms upon the re-alignment orrelaxation of the nuclear spin of atoms in an object being imaged (e.g.,atoms in the tissue of the human body). Detected MR signals may beprocessed to produce images, which in the context of medicalapplications, allows for the investigation of internal structures and/orbiological processes within the body for diagnostic, therapeutic and/orresearch purposes.

MRI provides an attractive imaging modality for biological imaging dueto the ability to produce non-invasive images having relatively highresolution and contrast without the safety concerns of other modalities(e.g., without needing to expose the subject to ionizing radiation,e.g., x-rays, or introducing radioactive material to the body).Additionally, MRI is particularly well suited to provide soft tissuecontrast, which can be exploited to image subject matter that otherimaging modalities are incapable of satisfactorily imaging. Moreover, MRtechniques are capable of capturing information about structures and/orbiological processes that other modalities are incapable of acquiring.However, there are a number of drawbacks to MRI that, for a givenimaging application, may involve the relatively high cost of theequipment, limited availability and/or difficulty in gaining access toclinical MRI scanners and/or the length of the image acquisitionprocess.

The trend in clinical MRI has been to increase the field strength of MRIscanners to improve one or more of scan time, image resolution, andimage contrast, which, in turn, continues to drive up costs. The vastmajority of installed MRI scanners operate at 1.5 or 3 tesla (T), whichrefers to the field strength of the main magnetic field B₀. A rough costestimate for a clinical MRI scanner is approximately one million dollarsper tesla, which does not factor in the substantial operation, service,and maintenance costs involved in operating such MRI scanners.

Additionally, conventional high-field MRI systems typically requirelarge superconducting magnets and associated electronics to generate astrong uniform static magnetic field (B₀) in which an object (e.g., apatient) is imaged. The size of such systems is considerable with atypical MRI installment including multiple rooms for the magnet,electronics, thermal management system, and control console areas. Thesize and expense of MRI systems generally limits their usage tofacilities, such as hospitals and academic research centers, which havesufficient space and resources to purchase and maintain them. The highcost and substantial space requirements of high-field MRI systemsresults in limited availability of MRI scanners. As such, there arefrequently clinical situations in which an MRI scan would be beneficial,but due to one or more of the limitations discussed above, is notpractical or is impossible, as discussed in further detail below.

SUMMARY

Some embodiments include a switching circuit configured to be coupled toa radio-frequency (RF) coil of a low-field magnetic resonance imagingsystem. The switching circuit comprises at least one field effecttransistor (FET) configured to operate as an RF switch at an operatingfrequency of less than 10 MHz.

Some embodiments include a drive circuit configured to apply a gatevoltage to at least one field-effect transistor (FET) configured tooperate as a radio-frequency switch in a low-field magnetic resonanceimaging system. The drive circuit comprises at least one isolationelement configured to isolate a voltage source from the at least oneFET.

Some embodiments include a circuit configured to tune a radio frequency(RF) coil coupled to an amplifier of a low-field magnetic resonanceimaging system. The circuit comprises tuning circuitry coupled acrossinputs of the amplifier, and active feedback circuitry coupled betweenan output of the amplifier and an input of the amplifier.

Some embodiments include a circuit configured to tune a radio frequency(RF) coil coupled to an amplifier of a low-field magnetic resonanceimaging system. The circuit comprises active feedback circuitry coupledbetween an output of the amplifier and an input of the amplifier toreduce a quality factor of the RF coil.

Some embodiments include a method of tuning a radio frequency (RF) coilcoupled to an amplifier of a low-field magnetic resonance imagingsystem. The method comprises arranging tuning circuitry across first andsecond inputs of the amplifier, and coupling active feedback circuitrybetween an output of the amplifier and an input of the amplifier.

Some embodiments include a radio-frequency (RF) coil for use in alow-field magnetic resonance imaging system. The RF coil comprises asubstrate having a first side and a second side, and a conductorincluding a first portion wound around the substrate from the first sideto the second side at a first plurality of locations spaced between thefirst side and the second side and a second portion wound around thesubstrate from the second side to the first side at a second pluralityof locations spaced between the first side and the second side, whereinthe first plurality of locations alternate with the second plurality oflocations spaced between the first side and the second side.

Some embodiments include a method of manufacturing a radio-frequency(RF) coil for use in a low-field magnetic resonance imaging system. Themethod comprises providing a substrate having circumferential groovesformed therein at a plurality of levels, each of which is arranged at adifferent distance from a first side of the substrate and connectinggrooves that connect adjacent levels of the plurality levels, windingwithin a first part of the circumferential grooves and the connectinggrooves, a first portion of a conductor from the first side of thesubstrate to a second side of the substrate, and winding within a secondpart of the circumferential grooves and the connecting grooves, a secondportion of the conductor from the second side of the substrate to thefirst side of the substrate, wherein the first part of thecircumferential grooves and the second part of the circumferentialgrooves do not overlap.

Some embodiments include a radio-frequency (RF) coil for use in alow-field magnetic resonance imaging system. The RF coil comprises asubstrate having a first side and a second side, and a conductor woundaround the substrate in a balanced winding pattern, wherein, in thebalanced winding pattern, a first portion of the conductor wound aroundthe substrate from the first side to the second side crosses over asecond portion of the conductor wound around the substrate from thesecond side to the first side.

The foregoing apparatus and method embodiments may be implemented withany suitable combination of aspects, features, and acts described aboveor in further detail below. These and other aspects, embodiments, andfeatures of the present teachings can be more fully understood from thefollowing description in conjunction with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

Various aspects and embodiments of the disclosed technology will bedescribed with reference to the following figures. It should beappreciated that the figures are not necessarily drawn to scale.

FIG. 1 illustrates exemplary components of a magnetic resonance imagingsystem;

FIG. 2 illustrates components of an RF signal chain for a magneticresonance imaging system;

FIG. 3 illustrates decoupling circuitry for use within an RF signalchain of a magnetic resonance imaging system;

FIGS. 4A and 4B show simulation data associated with the decouplingcircuitry of FIG. 3;

FIG. 5 illustrates decoupling circuitry using amplifier feedback inaccordance with some embodiments;

FIG. 6 illustrates the decoupling circuitry of FIG. 5 in which a singlecapacitor is used to provide coil tuning in accordance with someembodiments;

FIGS. 7A and 7B show simulation data associated with the decouplingcircuitry of FIG. 6;

FIG. 8 illustrates the decoupling circuitry of FIG. 5 in which atuning/matching network is used to provide coil tuning in accordancewith some embodiments;

FIG. 9 illustrates circuitry for a diode-based switch for use in an RFsignal chain of a magnetic resonance imaging system;

FIG. 10 illustrates circuitry for a FET-based switch for use in an RFsignal chain of a magnetic resonance imaging system in accordance withsome embodiments;

FIGS. 11A-C illustrate operating conditions of a FET in the FET-basedswitch of FIG. 10.

FIG. 12 illustrates circuitry for a GaN FET-based switch for use in anRF signal chain of a magnetic resonance imaging system in accordancewith some embodiments;

FIG. 13 illustrates circuitry for a GaN FET-based switch for use in anRF signal chain of a magnetic resonance imaging system in accordancewith some embodiments;

FIG. 14A illustrates a single-pass winding design for an RF coil;

FIG. 14B illustrates a top view of the single-pass winding design ofFIG. 14A;

FIG. 15A illustrates a schematic of an RF coil connected to an amplifierfor use in a low-field MRI system in accordance with some embodiments;

FIG. 15B illustrates an impedance model for representing impedancesassociation with a parasitic coupling of an object to an RF coil inaccordance with some embodiments;

FIG. 16A illustrates an interlaced winding design for an RF coil inaccordance with some embodiments;

FIG. 16B illustrates a top view of the interlaced winding design of FIG.16A;

FIG. 17A illustrates an alternate interlaced winding design for an RFcoil in accordance with some embodiments;

FIG. 17B illustrates a top view of the interlaced winding design of FIG.17A;

FIG. 18A illustrates a double loop winding design for an RF coil inaccordance with some embodiments;

FIG. 18B illustrates a top view of the double loop winding design ofFIG. 18A;

FIG. 19 illustrates a top view of a reverse helix winding design for anRF coil in accordance with some embodiments;

FIG. 20 illustrates a balanced winding design for an RF surface coil inaccordance with some embodiments;

FIG. 21 illustrates a process for creating an RF coil with a balancedwinding pattern in accordance with some embodiments;

FIGS. 22A-22L illustrate acts in a process for creating an RF coil withan interlaced winding pattern in accordance with some embodiments;

FIGS. 23A-23H illustrate acts in a process for creating an RF coil withan alternate interlaced winding pattern in accordance with someembodiments; and

FIG. 24 illustrates decoupling circuitry using mutual inductive feedbackin accordance with some embodiments.

DETAILED DESCRIPTION

The MRI scanner market is overwhelmingly dominated by high-fieldsystems, and particularly for medical or clinical MRI applications. Asdiscussed above, the general trend in medical imaging has been toproduce MRI scanners with increasingly greater field strengths, with thevast majority of clinical MRI scanners operating at 1.5 T or 3 T, withhigher field strengths of 7 T and 9 T used in research settings. As usedherein, “high-field” refers generally to MRI systems presently in use ina clinical setting and, more particularly, to MRI systems operating witha main magnetic field (i.e., a B₀ field) at or above 1.5 T, thoughclinical systems operating between 0.5 T and 1.5 T are often alsocharacterized as “high-field.” Field strengths between approximately 0.2T and 0.5 T have been characterized as “mid-field” and, as fieldstrengths in the high-field regime have continued to increase, fieldstrengths in the range between 0.5 T and 1 T have also beencharacterized as mid-field. By contrast, “low-field” refers generally toMRI systems operating with a B₀ field of less than or equal toapproximately 0.2 T, though systems having a B₀ field of between 0.2 Tand approximately 0.3 T have sometimes been characterized as low-fieldas a consequence of increased field strengths at the high end of thehigh-field regime. Within the low-field regime, low-field MRI systemsoperating with a B₀ field of less than 0.1 T are referred to herein as“very low-field” and low-field MRI systems operating with a B₀ field ofless than 10 mT are referred to herein as “ultra-low field.”

As discussed above, conventional MRI systems require specializedfacilities. An electromagnetically shielded room is required for the MRIsystem to operate and the floor of the room must be structurallyreinforced. Additional rooms must be provided for the high-powerelectronics and the scan technician's control area. Secure access to thesite must also be provided. In addition, a dedicated three-phaseelectrical connection must be installed to provide the power for theelectronics that, in turn, are cooled by a chilled water supply.Additional HVAC capacity typically must also be provided. These siterequirements are not only costly, but significantly limit the locationswhere MRI systems can be deployed. Conventional clinical MRI scannersalso require substantial expertise to both operate and maintain. Thesehighly trained technicians and service engineers add large on-goingoperational costs to operating an MRI system. Conventional MRI, as aresult, is frequently cost prohibitive and is severely limited inaccessibility, preventing MRI from being a widely available diagnostictool capable of delivering a wide range of clinical imaging solutionswherever and whenever needed. Typically, a patient must visit one of alimited number of facilities at a time and place scheduled in advance,preventing MRI from being used in numerous medical applications forwhich it is uniquely efficacious in assisting with diagnosis, surgery,patient monitoring and the like.

As discussed above, high-field MRI systems require specially adaptedfacilities to accommodate the size, weight, power consumption andshielding requirements of these systems. For example, a 1.5 T MRI systemtypically weighs between 4-10 tons and a 3 T MRI system typically weighsbetween 8-20 tons. In addition, high-field MRI systems generally requiresignificant amounts of heavy and expensive shielding. Many mid-fieldscanners are even heavier, weighing between 10-20 tons due, in part, tothe use of very large permanent magnets and/or yokes. Commerciallyavailable low-field MRI systems (e.g., operating with a B₀ magneticfield of 0.2 T) are also typically in the range of 10 tons or more duethe large of amounts of ferromagnetic material used to generate the B₀field, with additional tonnage in shielding. To accommodate this heavyequipment, rooms (which typically have a minimum size of 30-50 squaremeters) have to be built with reinforced flooring (e.g., concreteflooring), and must be specially shielded to prevent electromagneticradiation from interfering with operation of the MRI system. Thus,available clinical MRI systems are immobile and require the significantexpense of a large, dedicated space within a hospital or facility, andin addition to the considerable costs of preparing the space foroperation, require further additional on-going costs in expertise inoperating and maintaining the system.

In addition, currently available MRI systems typically consume largeamounts of power. For example, common 1.5 T and 3 T MRI systemstypically consume between 20-40 kW of power during operation, whileavailable 0.5 T and 0.2 T MRI systems commonly consume between 5-20 kW,each using dedicated and specialized power sources. Unless otherwisespecified, power consumption is referenced as average power consumedover an interval of interest. For example, the 20-40 kW referred toabove indicates the average power consumed by conventional MRI systemsduring the course of image acquisition, which may include relativelyshort periods of peak power consumption that significantly exceeds theaverage power consumption (e.g., when the gradient coils and/or RF coilsare pulsed over relatively short periods of the pulse sequence).Intervals of peak (or large) power consumption are typically addressedvia power storage elements (e.g., capacitors) of the MRI system itself.Thus, the average power consumption is the more relevant number as itgenerally determines the type of power connection needed to operate thedevice. As discussed above, available clinical MRI systems must havededicated power sources, typically requiring a dedicated three-phaseconnection to the grid to power the components of the MRI system.Additional electronics are then needed to convert the three-phase powerinto single-phase power utilized by the MRI system. The many physicalrequirements of deploying conventional clinical MRI systems creates asignificant problem of availability and severely restricts the clinicalapplications for which MRI can be utilized.

Accordingly, the many requirements of high-field MRI renderinstallations prohibitive in many situations, limiting their deploymentto large institutional hospitals or specialized facilities and generallyrestricting their use to tightly scheduled appointments, requiring thepatient to visit dedicated facilities at times scheduled in advance.Thus, the many restrictions on high field MRI prevent MRI from beingfully utilized as an imaging modality. Despite the drawbacks ofhigh-field MRI mentioned above, the appeal of the significant increasein SNR at higher fields continues to drive the industry to higher andhigher field strengths for use in clinical and medical MRI applications,further increasing the cost and complexity of MRI scanners, and furtherlimiting their availability and preventing their use as ageneral-purpose and/or generally-available imaging solution.

The low SNR of MR signals produced in the low-field regime (particularlyin the very low-field regime) has prevented the development of arelatively low cost, low power and/or portable MRI system. Conventional“low-field” MRI systems operate at the high end of what is typicallycharacterized as the low-field range (e.g., clinically availablelow-field systems have a floor of approximately 0.2 T) to achieve usefulimages. Though somewhat less expensive then high-field MRI systems,conventional low-field MRI systems share many of the same drawbacks. Inparticular, conventional low-field MRI systems are large, fixed andimmobile installments, consume substantial power (requiring dedicatedthree-phase power hook-ups) and require specially shielded rooms andlarge dedicated spaces. The challenges of low-field MRI have preventedthe development of relatively low cost, low power and/or portable MRIsystems that can produce useful images.

The inventors have developed techniques enabling portable, low-field,low power and/or lower-cost MRI systems that can improve the wide-scaledeployability of MRI technology in a variety of environments beyond thecurrent MRI installments at hospitals and research facilities. As aresult, MRI can be deployed in emergency rooms, small clinics, doctor'soffices, in mobile units, in the field, etc. and may be brought to thepatient (e.g., bedside) to perform a wide variety of imaging proceduresand protocols. Some embodiments include very low-field MRI systems(e.g., 0.1 T, 50 mT, 20 mT, etc.) that facilitate portable, low-cost,low-power MRI, significantly increasing the availability of MRI in aclinical setting.

There are numerous challenges to developing a clinical MRI system in thelow-field regime. As used herein, the term clinical MRI system refers toan MRI system that produces clinically useful images, which refers toimages having sufficient resolution and adequate acquisition times to beuseful to a physician or clinician for its intended purpose given aparticular imaging application. As such, the resolutions/acquisitiontimes of clinically useful images will depend on the purpose for whichthe images are being obtained. Among the numerous challenges inobtaining clinically useful images in the low-field regime is therelatively low SNR. Specifically, the relationship between SNR and B₀field strength is approximately B₀ ^(5/4) at field strength above 0.2 Tand approximately B₀ ^(3/2) at field strengths below 0.1 T. As such, theSNR drops substantially with decreases in field strength with even moresignificant drops in SNR experienced at very low field strength. Thissubstantial drop in SNR resulting from reducing the field strength is asignificant factor that has prevented development of clinical MRIsystems in the very low-field regime. In particular, the challenge ofthe low SNR at very low field strengths has prevented the development ofa clinical MRI system operating in the very low-field regime. As aresult, clinical MRI systems that seek to operate at lower fieldstrengths have conventionally achieved field strengths of approximatelythe 0.2 T range and above. These MRI systems are still large, heavy andcostly, generally requiring fixed dedicated spaces (or shielded tents)and dedicated power sources.

The inventors have developed low-field and very low-field MRI systemscapable of producing clinically useful images, allowing for thedevelopment of portable, low cost and easy to use MRI systems notachievable using state of the art technology. According to someembodiments, an MRI system can be transported to the patient to providea wide variety of diagnostic, surgical, monitoring and/or therapeuticprocedures, generally, whenever and wherever needed.

FIG. 1 is a block diagram of typical components of a MRI system 100. Inthe illustrative example of FIG. 1, MRI system 100 comprises computingdevice 104, controller 106, pulse sequences store 108, power managementsystem 110, and magnetics components 120. It should be appreciated thatsystem 100 is illustrative and that a MRI system may have one or moreother components of any suitable type in addition to or instead of thecomponents illustrated in FIG. 1. However, a MRI system will generallyinclude these high level components, though the implementation of thesecomponents for a particular MRI system may differ vastly, as discussedin further detail below.

As illustrated in FIG. 1, magnetics components 120 comprise B₀ magnet122, shim coils 124, RF transmit and receive coils 126, and gradientcoils 128. Magnet 122 may be used to generate the main magnetic fieldB₀. Magnet 122 may be any suitable type or combination of magneticscomponents that can generate a desired main magnetic B₀ field. Asdiscussed above, in the high field regime, the B₀ magnet is typicallyformed using superconducting material generally provided in a solenoidgeometry, requiring cryogenic cooling systems to keep the B₀ magnet in asuperconducting state. Thus, high-field B₀ magnets are expensive,complicated and consume large amounts of power (e.g., cryogenic coolingsystems require significant power to maintain the extremely lowtemperatures needed to keep the B₀ magnet in a superconducting state),require large dedicated spaces, and specialized, dedicated powerconnections (e.g., a dedicated three-phase power connection to the powergrid). Conventional low-field B₀ magnets (e.g., B₀ magnets operating at0.2 T) are also often implemented using superconducting material andtherefore have these same general requirements. Other conventionallow-field B₀ magnets are implemented using permanent magnets, which toproduce the field strengths to which conventional low-field systems arelimited (e.g., between 0.2 T and 0.3 T due to the inability to acquireuseful images at lower field strengths), need to be very large magnetsweighing 5-20 tons. Thus, the B₀ magnet of conventional MRI systemsalone prevents both portability and affordability.

Gradient coils 128 may be arranged to provide gradient fields and, forexample, may be arranged to generate gradients in the B₀ field in threesubstantially orthogonal directions (X, Y, Z). Gradient coils 128 may beconfigured to encode emitted MR signals by systematically varying the B₀field (the B₀ field generated by magnet 122 and/or shim coils 124) toencode the spatial location of received MR signals as a function offrequency or phase. For example, gradient coils 128 may be configured tovary frequency or phase as a linear function of spatial location along aparticular direction, although more complex spatial encoding profilesmay also be provided by using nonlinear gradient coils. For example, afirst gradient coil may be configured to selectively vary the B₀ fieldin a first (X) direction to perform frequency encoding in thatdirection, a second gradient coil may be configured to selectively varythe B₀ field in a second (Y) direction substantially orthogonal to thefirst direction to perform phase encoding, and a third gradient coil maybe configured to selectively vary the B₀ field in a third (Z) directionsubstantially orthogonal to the first and second directions to enableslice selection for volumetric imaging applications. As discussed above,conventional gradient coils also consume significant power, typicallyoperated by large, expensive gradient power sources, as discussed infurther detail below.

MRI is performed by exciting and detecting emitted MR signals usingtransmit and receive coils, respectively (often referred to as radiofrequency (RF) coils). Transmit/receive coils may include separate coilsfor transmitting and receiving, multiple coils for transmitting and/orreceiving, or the same coils for transmitting and receiving. Thus, atransmit/receive component may include one or more coils fortransmitting, one or more coils for receiving and/or one or more coilsfor transmitting and receiving. Transmit/receive coils are also oftenreferred to as Tx/Rx or Tx/Rx coils to generically refer to the variousconfigurations for the transmit and receive magnetics component of anMRI system. These terms are used interchangeably herein. In FIG. 1, RFtransmit and receive coils 126 comprise one or more transmit coils thatmay be used to generate RF pulses to induce an oscillating magneticfield B₁. The transmit coil(s) may be configured to generate anysuitable types of RF pulses.

Power management system 110 includes electronics to provide operatingpower to one or more components of the low-field MRI system 100. Forexample, as discussed in more detail below, power management system 110may include one or more power supplies, gradient power components,transmit coil components, and/or any other suitable power electronicsneeded to provide suitable operating power to energize and operatecomponents of MRI system 100. As illustrated in FIG. 1, power managementsystem 110 comprises power supply 112, power component(s) 114,transmit/receive switch 116, and thermal management components 118(e.g., cryogenic cooling equipment for superconducting magnets). Powersupply 112 includes electronics to provide operating power to magneticcomponents 120 of the MRI system 100. For example, power supply 112 mayinclude electronics to provide operating power to one or more B₀ coils(e.g., B₀ magnet 122) to produce the main magnetic field for thelow-field MRI system. Transmit/receive switch 116 may be used to selectwhether RF transmit coils or RF receive coils are being operated.

Power component(s) 114 may include one or more RF receive (Rx)pre-amplifiers that amplify MR signals detected by one or more RFreceive coils (e.g., coils 126), one or more RF transmit (Tx) powercomponents configured to provide power to one or more RF transmit coils(e.g., coils 126), one or more gradient power components configured toprovide power to one or more gradient coils (e.g., gradient coils 128),and one or more shim power components configured to provide power to oneor more shim coils (e.g., shim coils 124).

In conventional MRI systems, the power components are large, expensiveand consume significant power. Typically, the power electronics occupy aroom separate from the MRI scanner itself. The power electronics notonly require substantial space, but are expensive complex devices thatconsume substantial power and require wall mounted racks to besupported. Thus, the power electronics of conventional MRI systems alsoprevent portability and affordable of MRI.

As illustrated in FIG. 1, MRI system 100 includes controller 106 (alsoreferred to as a console) having control electronics to sendinstructions to and receive information from power management system110. Controller 106 may be configured to implement one or more pulsesequences, which are used to determine the instructions sent to powermanagement system 110 to operate the magnetic components 120 in adesired sequence (e.g., parameters for operating the RF transmit andreceive coils 126, parameters for operating gradient coils 128, etc.).As illustrated in FIG. 1, controller 106 also interacts with computingdevice 104 programmed to process received MR data. For example,computing device 104 may process received MR data to generate one ormore MR images using any suitable image reconstruction process(es).Controller 106 may provide information about one or more pulse sequencesto computing device 104 for the processing of data by the computingdevice. For example, controller 106 may provide information about one ormore pulse sequences to computing device 104 and the computing devicemay perform an image reconstruction process based, at least in part, onthe provided information. In conventional MRI systems, computing device104 typically includes one or more high performance work-stationsconfigured to perform computationally expensive processing on MR datarelatively rapidly. Such computing devices are relatively expensiveequipment on their own.

As should be appreciated from the foregoing, currently availableclinical MRI systems (including high-field, mid-field and low-fieldsystems) are large, expensive, fixed installations requiring substantialdedicated and specially designed spaces, as well as dedicated powerconnections. The inventors have developed low-field, including very-lowfield, MRI systems that are lower cost, lower power and/or portable,significantly increasing the availability and applicability of MRI.According to some embodiments, a portable MRI system is provided,allowing an MRI system to be brought to the patient and utilized atlocations where it is needed.

As discussed above, some embodiments include an MRI system that isportable, allowing the MRI device to be moved to locations in which itis needed (e.g., emergency and operating rooms, primary care offices,neonatal intensive care units, specialty departments, emergency andmobile transport vehicles and in the field). There are numerouschallenges that face the development of a portable MRI system, includingsize, weight, power consumption and the ability to operate in relativelyuncontrolled electromagnetic noise environments (e.g., outside aspecially shielded room.

An aspect of portability involves the capability of operating the MRIsystem in a wide variety of locations and environments. As discussedabove, currently available clinical MRI scanners are required to belocated in specially shielded rooms to allow for correct operation ofthe device and is one (among many) of the reasons contributing to thecost, lack of availability and non-portability of currently availableclinical MRI scanners. Thus, to operate outside of a specially shieldedroom and, more particularly, to allow for generally portable, cartableor otherwise transportable MRI, the MRI system must be capable ofoperation in a variety of noise environments. The inventors havedeveloped noise suppression techniques that allow the MRI system to beoperated outside of specially shielded rooms, facilitating bothportable/transportable MRI as well as fixed MRI installments that do notrequire specially shielded rooms. While the noise suppression techniquesallow for operation outside specially shielded rooms, these techniquescan also be used to perform noise suppression in shielded environments,for example, less expensive, loosely or ad-hoc shielding environments,and can be therefore used in conjunction with an area that has beenfitted with limited shielding, as the aspects are not limited in thisrespect.

A further aspect of portability involves the power consumption of theMRI system. As also discussed above, current clinical MRI systemsconsume large amounts of power (e.g., ranging from 20 kW to 40 kWaverage power consumption during operation), thus requiring dedicatedpower connections (e.g., dedicated three-phase power connections to thegrid capable of delivering the required power). The requirement of adedicated power connection is a further obstacle to operating an MRIsystem in a variety of locations other than expensive dedicated roomsspecially fitted with the appropriate power connections. The inventorshave developed low power MRI systems capable of operating using mainselectricity such as a standard wall outlet (e.g., 120V/20A connection inthe U.S.) or common large appliance outlets (e.g., 220-240V/30A),allowing the device to be operated anywhere common power outlets areprovided. The ability to “plug into the wall” facilitates bothportable/transportable MRI as well as fixed MRI system installationswithout requiring special, dedicated power such as a three-phase powerconnection.

As discussed above, a portable MRI device designed in accordance withthe techniques described herein includes RF transmit and receive coils126 configured to generate a B₁ magnetic field during a transmitoperation and to collect flux from an MR signal generated by an imagedobject during a receive operation. Signals sensed by the RF receive coilare amplified and processed prior to conversion into MR images.Circuitry involved in the control and processing of signals recorded bythe RF receive coils 126 are referred to herein as “RF signal chain”circuitry. The inventors have recognized that components of the RFsignal chain circuitry used in conventional high-field MRI system arenot appropriate and/or optimized for use in a low-field MRI systemdesigned in accordance with the techniques described herein. To thisend, some embodiments are directed to improved RF signal chain circuitryfor use in a portable low-field MRI system.

FIG. 2 schematically illustrates some components of RF signal chaincircuitry 200 included in some embodiments. RF signal chain circuitry200 includes RF transmit/receive coil 210 and transmit/receive circuitry212 configured to selectively couple the RF coil 210 to the RF receivecircuitry depending on whether the RF coil 210 is being operated totransmit or receive. To operate optimally, RF coils are often tuned toresonate as close as possible to a particular frequency called theLarmor frequency. The Larmor frequency (w) is related to the strength ofthe B₀ field in accordance with following relation: ω=γB, where γ is thegyromagnetic ratio of the imaged isotope (e.g., 1H) in MHz/T, and B isthe strength of the B₀ field in Tesla. Examples of commonly used Larmorfrequencies used in high-field MRI are approximately 64 MHz for a 1.5 TMRI system and approximately 128 MHz for a 3 T MRI system. For low-fieldMRI systems, the Larmor frequency is substantially lower than forhigh-field MRI systems. For example, the Larmor frequency for a 64 mTMRI system is approximately 2.75 MHz. RF signal chain circuitry 200further includes tuning/matching circuitry 214 configured to transformthe impedance of RF coil 210 to optimize performance. The output oftuning/matching circuitry 214 is provided to amplifier 216 (e.g., a lownoise amplifier), which amplifies the RF signals prior to conversioninto image signals. The inventors have recognized one of thedifficulties with using RF coils in low-field MRI systems is thesusceptibility of such coils to noise in electronic components. Inaccordance with some embodiments, one or more of components 210, 214,and 216 are configured to reduce noise in the RF signal chain.

Some embodiments include multiple RF coils to improve thesignal-to-noise ratio (SNR) of signals detected by an RF coil network.For example, a collection of RF coils may be arranged at differentlocations and orientations to detect a comprehensive RF field. Accordingto some embodiments, a portable MRI system comprises multiple RFtransmit/receive coils to improve the SNR of image acquisition. Forexample, a portable MRI system may comprise 2, 4, 8, 16, 32 or more RFreceive coils to improve the SNR of MR signal detection.

As discussed above, in general, RF coils are tuned to increase coilsensitivity at a frequency of interest (e.g., the Larmor frequency).However, inductive coupling between adjacent or neighboring coils (e.g.,RF coils sufficiently proximate one another) degrades the sensitivity oftuned coils and significantly reduces the effectiveness of thecollection of the RF coils. Techniques for geometrically decouplingneighboring coils exist but place strict constraints on coil orientationand position in space, reducing the ability of the collection of RFcoils to accurately detect the RF field and, as a consequence, degradingthe signal-to-noise performance.

To address the negative impact of inductive coupling between coils, theinventors have utilized coil decoupling techniques that reduce theeffect of inductive coupling between radio frequency coils in multi-coiltransmit/receive systems. For example, FIG. 3 illustrates a passivedecoupling circuit 300 configured to reduce inductive coupling betweenradio frequency coils in a multi-coil transmit/receive system. Circuit300 is configured to decouple RF coils that may be subjected to B₁transmit fields (e.g., from an RF transmit coil). The purpose of thedecoupling circuit is to minimize the current through the RF coil for agiven AC excitation voltage at the Larmor frequency. In particular,inductor L1 represents an RF signal coil within the field of view of theMRI system. Capacitors C1 and C2 form a tuning circuit that matches theinductance of the coil to the input of the low noise amplifier (LNA) tooptimize noise performance impedance. Inductor L2 and capacitor C3 forma tank circuit that reduces the current that flows in the loop thatincludes L1, C1 and C2 to prevent coupling of the RF coil to othercoils. FIG. 4A illustrates a plot of the voltage at the LNA input at theresonant frequency of the RF coil based on a simulation of circuit 300in FIG. 3. FIG. 4B illustrates a plot of current through the RF coilbased on a simulation of circuit 300 in FIG. 3. As shown, at theresonant frequency of 2.75 MHz, the LNA voltage is about 26 dB (FIG. 4A)and the coil current is −37 dB (FIG. 4B). In each of FIGS. 4A and 4B,magnitude of the measured quantity is represented as a solid line andphase of the measured quantity is represented as a dashed line.

The inventors have recognized that decoupling using a tuned matchingfilter to reduce the current in the RF coil has some drawbacks includingthe need to tune multiple components (e.g., capacitors C1, C2 and C3) tothe operating frequency of the coil. Additionally, losses in theinductor L2 result in a loss of SNR. As such, decoupling efficiency is atrade off with SNR efficiency. Furthermore, as shown in FIG. 4B,although the tuned matching filter reduces the coil currentsubstantially at the resonant frequency, the sharp valley in the currentwaveform demonstrates that the current reduction through the RF coil isonly small for a limited bandwidth surrounding the resonant frequency.

Some embodiments are directed to an improved decoupling circuitconfigured to reduce the current in the RF coil by damping the coilresponse using feedback from the output of the amplifier. FIG. 5 showsan example of a decoupling circuit 400 configured to provide feedbackdecoupling in accordance with some embodiments. Circuit 400 includes anactive feedback path from the output of an amplifier LNA to an input ofthe LNA. In the example shown in FIG. 5, the active feedback pathincludes a single feedback path. However, it should be appreciated thatthe active feedback path may alternatively be implemented as a pluralityof feedback paths, each which provides a different type of feedbackdecoupling when selected. For example, in some embodiments, the activefeedback path includes a first feedback path configured to provide afirst feedback signal and a second feedback path configured to provide asecond feedback signal.

The inventors have recognized that the phase of the feedback signalaffects the amplification gain at the tuning frequency. For example, insome embodiments that include multiple feedback paths in the activefeedback path, a first feedback path may provide a first feedback signal90 or 270 degrees out of phase with a resonant frequency of the RF coiland a second feedback path may provide a second feedback signal 180degrees out of phase with a resonant frequency of the RF coil.Alternatively, the gain of the amplifier may be tuned to be 90 or 270degrees out of phase with the resonant frequency of the coil. When aphase of 270 degrees is used, the amplification gain at the tuningfrequency may be maximum. In other embodiments in which a singlefeedback path is used, the phase of the feedback signal may be set to180 degrees to provide more efficient decoupling due to less current inthe coil.

The feedback decoupling provided by circuit 400 uses active negativefeedback to damp the coil response (also referred to as reducing thequality (Q) factor of the coil or “de-Qing” the coil) and therebyreducing current flowing in the RF coil. As shown, circuit 400 alsoincludes a tuning/matching circuit arranged between the RF coil and theLNA. Any suitable tuning/matching circuit may be used in accordance withsome embodiments, examples of which are described below.

FIG. 6 illustrates a feedback-based decoupling circuit 500 thatimplements the tuning/matching circuit using a single capacitor C1. Incontrast to decoupling circuit 300, decoupling circuit 500 includes onlya single component (i.e., capacitor C1) to tune. Additionally, becausecircuit 500 only includes reactive components C1 and C2 and does notinclude an inductor in the tuning/matching circuit, the decouplingcircuit does not introduce the SNR losses associated with circuit 300due to the inclusion of inductor L1 in circuit 300.

Capacitor C1 may be implemented using a capacitor with a fixed value.Alternatively, capacitor C1 may be implemented using a capacitor with avariable value (e.g., a varactor diode). In yet further embodiments,capacitor C1 may be implemented using a capacitor with fixed value(e.g., 300 pF) arranged in parallel with a capacitor with variablevalue. Such an arrangement reduces the effect of AC losses introduced byuse of a variable capacitor in the feedback loop.

FIG. 7A illustrates a plot of the voltage at the LNA input at theresonant frequency of the RF coil based on a simulation of circuit 500in FIG. 5. FIG. 7B illustrates a plot of current through the RF coilbased on a simulation of circuit 500 in FIG. 6. As shown, at theresonant frequency of 2.75 MHz, the LNA input voltage is about 8 dB(FIG. 7B) and the current through the coil is about −35 dB. However, incontrast to the coil current plot shown in FIG. 4B which shows a sharpvalley at the resonant frequency of the coil, FIG. 7B shows that thecoil current is reduced over a much wider bandwidth when usingdecoupling circuit 500 as compared to circuit 300. Accordingly, incomparison to circuit 300, circuit 500 provides RF coil decoupling overa wider bandwidth.

FIG. 8 illustrates an alternative feedback-based decoupling circuit 600in which the single capacitor tuning/matching circuit of circuit 500shown in FIG. 6 is replaced with a tuning/matching network that includescomponents C1, C3 and L2. In circuit 600, a tuning/matching network isused to tune the RF coil (represented as L1) in addition to havingfeedback-based decoupling provided by an active feedback path includingcapacitor C2.

In some embodiments, the capacitive feedback circuitry provided, forexample, by the feedback components of circuits 400, 500, and 600 inFIGS. 5, 6 and 8, respectively, is replaced with mutual inductivefeedback circuitry. FIG. 24 illustrates an alternative feedback-baseddecoupling circuit 2400 in which the capacitive feedback circuitryshown, for example, in circuit 500 of FIG. 6 is replaced with mutualinductive feedback circuitry that includes components R1, R2 and L2. Incircuit 2400, inductors L1 and L2 are coupled mutually, e.g., using atransformer or through the air.

Another technique for providing RF coil decoupling in accordance withsome embodiments it to provide a transmit/receive switch in the RFsignal chain. The transmit/receive switch is configured to isolate theRF coil from the amplifier when RF signals are being transmitted by oneor more RF transmit coils. Specifically, the transmit/receive switchdivides the tuning/matching network into two network portions to protectsensitive electronics during RF transmit cycles. In some conventionalMRI systems (e.g., high-field MRI systems), the transmit/receive switch312 is typically implemented using a diode, such as a PIN diode. Anexample of transmit/receive switch circuitry that includes a diode D1 isshown in FIG. 9 as circuit 700. During a transmit pulse, diode D1 isturned on to create a short circuit, isolating the RF signal coil fromthe receive electronics. As described above in connection with circuit300, the resulting network provides a tank circuit with a high impedancethat ensures that the current in the RF coil remains small. Duringreceive cycles, diode D1 is turned off resulting in the RF coil beingconnected to the amplifier and being tuned by the tank circuitconfigured to limit the current through the RF coil, while allowing forsufficient signal to be detected at the output of the amplifier. Thus,the RF coil is connected to a first tank circuit during transmit cyclesand a second tank circuit during receive cycles of a pulse sequence.

Conventional decoupling circuits, such as that shown in FIG. 9, oftenuse PIN diodes to isolate the receive electronics from the RF signalcoil. However, PIN diodes suitable for performing this function in adecoupling circuit require approximately 0.1 A of current to turn thediode on. As an example, a transmit/receive coil system having eightreceive coils may require on the order of 0.8 A of current to decouplethe receive coils from the RF signal coil(s) for each transmit andreceive cycle of an image acquisition pulse sequence. Accordingly, overthe span of an image acquisition protocol, substantial power is consumedby the decoupling circuits of the RF transmit/receive system.Additionally, when PIN diodes are used, a biasing resistor R1 and an ACblocking filter including components L1 and C1 are required, and theground of the circuit is not isolated when the diode is in the offstate. Furthermore, while PIN diodes work well at higher frequenciesused in high-field MRI systems, PIN diodes do not work well at lowoperating frequencies (e.g., less than 10 MHz) used in low field or verylow field MRI systems. At such low frequencies, the PIN diode rectifiesthe signal rather than blocking it. For example, a DC bias currentI_(bias) allows the diode to be forward biased even when a negativesignal is applied. For an AC signal of frequency f and peak currentI_(peak), the ratio I_(peak)/f needs to be lower than the product of theDC bias current I_(bias) and the carrier lifetime τ in accordance withthe following relation: τI_(bias)>I_(peak)/f for the PIN diode tofunction properly to block the signal. However, some low-field MRIapplications may have the following parameters, I_(peak)=10A, f−2.75MHz, I_(bias)=100 mA. According to the relation above, for theseparameters, the PIN diode would need to have a carrier lifetime τ>37 μs,which is not a characteristics of commercially-available PIN diodes.

The inventors have recognized that PIN diodes typically used in adecoupling circuit may be replaced by Gallium Nitride (GaN) field effecttransistors (FETs) to address some of the shortcomings of using PINdiodes in an RF transmit/receive circuit of a low-field MRI systemincluding reducing the power consumption of the RF transmit/receivesystem. In particular, GaN FETs require on the order of microamps toturn on, reducing the power consumption by several orders of magnitude.In addition, the resistance of the GaN FETs when turned on is smallcompared to PIN diodes, reducing negative impact on the tank circuit.According to some embodiments, diode D1 in circuit 700 is replaced withone or more GaN FETs, thereby reducing the power consumption of the RFtransmit/receive system.

FIG. 10 illustrates a RF transmit/receive switch circuit 412 inaccordance with some embodiments, in which the diode D1 of circuit 700has been replaced with a pair of mirrored FETs (e.g., GaN FETs), F1 andF2. Although circuit 412 includes a pair of mirrored FETs, in someembodiments, an RF transmit/receive switch circuit 412 may include anysuitable number of FETs including, but not limited to, a single FET.Unlike PIN diodes, GaN FETs operate well at all frequencies, havenegligible power consumption, are ground isolated in the off-state, andhave a lower on-state resistance (e.g., <0.1 Ohm) than PIN diodes.

FIGS. 11A-C illustrate operating states of a FET used as a switch in anRF transmit/receive system in accordance with some embodiments. FIG. 11Aillustrates a GaN FET configured as a switch between a drain node D anda source node S. The gate G of the GaN FET is used to control the stateof the switch between on and off. FIG. 11B illustrates that in the offstate, the GaN FET can be modeled with three lumped capacitors C_ds,C_gs, and C_gd. In such a configuration, the drain D is isolated fromthe source S provided that the value of C_ds is small (e.g., 10-100 pF).In some embodiments, the drain-source capacitance of at least one GaNFET included in a transmit/receive switch is at least 15 pF. FIG. 11Cillustrates that in the on state, the drain-source capacitance C_ds isreplaced by a short circuit.

FIG. 12 illustrates a circuit 1000 for driving a gate voltage on GaNFETs U1 and U2 arranged to operate as an RF transmit/receive switch inaccordance with some embodiments. The GaN FETs are configured to coupleand decouple the receive electronics from the RF coil. As shown,inductors L5 and L6 are arranged as a transformer configured to couple acontrol signal V2 to the gates of the FETs U1 and U2, while providingground isolation. The diode D3 operates to rectify the control signal tocreate a DC on/off voltage at the gates across capacitor C7. Theresistor R11 is configured to discharge capacitor C7 and the gatecapacitance of the FETs. The time constant of R11 and C7+Cgatesdetermines how quickly the transmit/receive switch turns off. In someembodiments, control signal V2 may be a 10 MHz sine wave coupled toinductor L5 to drive the FETs. In operation, the 10 MHz signal may beturned on/off to charge up the FET gates and then turned off. Thenresistor R11 discharges the gate drive to open the switch. In theexample of FIG. 12, the coupling between inductors L5 and L6 may be poorand the inductances small. For example, L5/L6 may be implemented in someembodiments as small air-core transformer or as an RF transformer.

FIG. 13 illustrates a drive circuit 1100 for driving a gate voltage onGaN FETs U1 and U2 arranged to operate as an RF transmit/receive switchin accordance with some embodiments. In circuit 1100 rather than usingan externally provided control signal V2 (as in circuit 1000), the RFtransmit pulse itself is used as the control signal to gate thetransmit/receive switch. In the example of FIG. 13, a coil representedby inductor L6 is configured to receive the RF transmit pulse, and inresponse, generate a voltage that drives the gates of the GaN FETs. Insome embodiments, each of the RF coils in an RF coil array may beassociated with a coil L6 configured to receive the RF transmit pulsefor that coil. In other embodiments, a subset (e.g., one) of the RFcoils in a multi-coil array may be associated with coil L6 configured toreceive the RF transmit pulse, and the switch signal generated by thecoil L6 in response to receiving the transmit pulse may be distributedto the circuitry associated with the other RF coils in the array.Circuit 1100 uses less complex drive circuitry than circuit 1000 becausea separate control signal generator is not required. However, aconsequence of using the transmit pulse as a control signal is that theswitch does not close until slightly after RF transmission begins andthe RF transmit pulse can be detected by pulse receiver coil L6.

Some embodiments are related to a novel design for a radio-frequency(RF) coil for use in a low-field MRI system. Some conventional RF coildesigns for use in MRI systems are configured as a solenoid, which wrapsaround an object to be imaged in a helix pattern. For example, headcoils commonly used in MRI systems include a conductor formed in asolenoid configuration such that a head of a person can be insertedinside of the solenoid. FIG. 14A schematically illustrates a solenoid RFcoil design in which a conductor is wound in a plurality of loops arounda substrate in a single pass from a first side of the substrate to asecond side of the substrate opposite to the first side. When the secondside of the substrate is reached, the conductor may be returned to thefirst side without forming additional loops, as shown. FIG. 14B shows atop view of the coil arrangement of FIG. 14A in which the loops ofconductor are represented by vertical lines. Points V+ and V− representthe ends of the conductor in the coil which in an MRI system areconnected to an amplifier (e.g., a low noise amplifier) configured toamplify the recorded signals.

In an ideal case, the potential recorded at the outputs of the RF coilare balanced such that V+−V−=0 in the absence of electromotive force(emf) in the coil. However, when an object, such as the head of aperson, is inserted into the solenoid coil, parasitic coupling occursbetween the object and the conductor in the coil that may result in V+and V− being unbalanced and producing a voltage at the amplifier input.The voltage is manifested as a noise signal in the recorded MR signalwhen the coil is used in the MRI system. Depending on the location ofthe head within the RF coil, the parasitic coupling may affect thesignals recorded at the points V+ and V− differently. For example, whenthe object is inserted at one end of the coil, the magnitude of thenoise introduced into the recorded signal due to parasitic coupling maybe larger at the point V+ compared to the point V− because of theshorter conductor distance between V+ and the point at which the noisewas introduced in the coil. Alternatively, if the object is arranged ator near the center of the coil between points V+ and V−, the noiseintroduced into the coil would affect the voltage detected at bothpoints V+ and V− equally. In yet another implementation, if the objectwas arranged closer to point V−, more noise would be detected at thepoint V− than the point V+ resulting in an unbalanced output (i.e.,V+−V−≠0).

FIG. 15A schematically illustrates that when an object (represented asvoltage source V₀) is inserted into a solenoid coil at a particularlocation, a parasitic coupling (represented as impedance Z_(C)) isintroduced into the coil at a single point in the coil. It should beappreciated that in practice, the parasitic coupling from the object tothe coil winding will be distributed. FIG. 15B illustrates an impedancemodel for how the introduction of the parasitic coupling affects thevoltages V₊, V⁻ measured at the ends of the conductor. Z_(C) representsthe parasitic coupling between the object and the coil, Z₊ representsthe impedance in the conductor between the point at which the parasiticcoupling is introduced and the point V₊, Z⁻ represents the impedance inthe conductor between the point at which the parasitic coupling isintroduced and the point V⁻, and Z_(G) represents the impedance betweeneach of the ends of the conductor (i.e., V₊ and V⁻) and ground. Whenthere is a weak parasitic coupling between the object and the coil(e.g., Z_(c) □ Z₊, Z⁻, Z_(G)), the follow relation describes thedifference in potential at the two ends of the conductor V₊ and V⁻:

${V_{+} - V_{-}} = {\frac{Z_{G}}{Z_{C}^{2}}\left( {Z_{-} - Z_{+}} \right)V_{0}}$

Because the outputs of the coil at V+ and V− may be unbalanced, someconventional RF coils include a balun between the RF coil and theamplifier to provide a balanced output and to reject common mode noiseintroduced into the coil. The inventors have recognized that the use ofbaluns to reject common mode noise introduced into an RF coil is notdesirable in a low-field MRI system due to small magnitude signals thatare received by the coil and the lossy characteristics of baluns. Tothis end, some embodiments are directed to an RF coil design that uses awinding pattern designed to reduce common mode noise, which mitigatesthe need to use a balun.

FIGS. 16-19 schematically illustrate RF coil designs in accordance withsome embodiments. For the coil designs shown in FIGS. 16-19, the resultis a solenoid coil having similar magnetic properties as theconventional coil winding pattern shown in FIG. 14A. For example, turnsof the coil located near each other detect a similar magnetic flux.However, the electrical properties of the RF coil designs shown in FIGS.16-19 are different than those for the FIG. 14A coil design. Inparticular, the winding designs shown in FIGS. 16-19 result in improvedbalance and common mode rejection when an object to be imaged isinserted in the coil. As shown, when an object to be imaged is insertedinto the coil, parasitic coupling between the object and the conductorresults in a voltage being induced in the turns of the conductor locatednear the object. For the winding patterns shown in FIGS. 16-19, adjacentturns of the conductor have a similar inductance and potential when avoltage is applied to the conductor because the adjacent turns arelocated a similar distance from the respective points V+ and V−.Accordingly, the voltage induced in the conductor due to parasiticcoupling of an object inserted in the coil to the conductor manifests asa noise signal similarly at both outputs V+ and V− regardless of thelocation of the object in the coil, such that V+−V−˜0, thereby reducingthe common mode noise.

Rather than winding the conductor in a single pass of loops from one endof the RF to the other end as shown in FIG. 14A, the conductor is woundin some embodiments in a balanced pattern using multiple (e.g., two)passes of loops or partial loops from end-to-end. FIG. 16A shows an“interlaced” winding pattern in which a conductor starting at one end ofa substrate is wound around the substrate by skipping portions of thesubstrate at different levels spaced from a first end of the conductoralong the winding direction in a first pass. When the conductor is woundfrom the other (second) end of the substrate in a second pass, theconductor is wound around those portions of the substrate that wereskipped in the first pass. FIG. 16B shows a top view of the interlacedwinding pattern illustrated in FIG. 16A.

The interlaced winding pattern shown in FIG. 16A skips complete turns(e.g., 360° revolutions) in a first pass and fills in those skippedturns in the second pass in the opposite direction. However, it shouldbe appreciated that alternate interlaced winding patterns are alsocontemplated. For example, FIG. 17A shows an interlaced pattern whererather than skipping a complete turn (e.g., 360°) in the first pass, thewinding pattern completes a series of half turns (e.g., 180°) around thesubstrate at each of a plurality of levels from the first end of thesubstrate while skipping the other half of the turns in the first pass.The skipped half turns are then filled in during the second pass fromthe second end to the first end of the substrate resulting in a fullsolenoid coil configuration. FIG. 17B shows a top view of the half turnskipping interlaced winding design of FIG. 17A. FIGS. 22A-L described inmore detail below illustrate a process for implementing the half turnskipping design illustrated in FIG. 17A.

FIG. 18A illustrates an example of an alternate balanced winding patternin accordance with some embodiments. In the winding pattern shown inFIG. 18A, a first plurality of loops of conductor are wound around thesubstrate from the first end to the second end without skipping anylevels on the first pass. On the second pass from the second end to thefirst end, a second plurality of loops located near the first pluralityof loops are wound around the substrate to create a “double” windingpattern. FIG. 18B shows a top view of the winding pattern illustrated inFIG. 18A.

FIG. 19 shows a top view of another balanced winding pattern with aninterlaced configuration in accordance with some embodiments. In thewinding pattern shown in FIG. 19, rather than forming loops in a seriesof levels from a first end of the substrate to a second end of thesubstrate (e.g., as shown in FIG. 16A), the conductor is wound in afirst pass from the first end to the second end of the substrate in ahelix configuration, and in a second pass from the second end to thefirst end of the substrate in a reverse helix configuration. Theparticular angle to form the helix winding configuration is not alimitation of embodiments of the invention, as any suitable angle withany desired number of turns around the substrate may be used.

The balanced winding patterns described above in connection with FIGS.16-19 relate to RF coils having a solenoid configuration in which anobject (e.g., a patient's head is inserted within the solenoid. Theinventors have recognized that the balanced winding techniques describedherein may also be used for coil configurations other than a solenoidcoil. For example, FIG. 20 illustrates an example of using a balancedwinding pattern to create an RF surface coil. The surface coil includestwo conductor windings arranged in close proximity to each other. Asshown, the distance (h) between the two windings may be made small(e.g., approaching a distance close to 0) such that the multiplewindings have magnetic properties similar to a coil that has a singlewinding. In some embodiments, the two windings may be configured to be180° out of phase with each other.

FIG. 21 illustrates a process 2100 for manufacturing an RF coil inaccordance with some embodiments. In act 2110, a substrate is providedaround which a conductor is to be wound. The substrate may be made ofany suitable non-magnetic material. In some embodiments the substratecomprises a plastic material fabricated, for example, using an additivemanufacturing process (e.g., 3D printing). Process 2130 then proceeds toact 2112, where a plurality of grooves are formed in the substrate. Forexample, the substrate may include a top and a bottom and the pluralityof grooves may be formed at locations space from the top of thesubstrate to the bottom of the substrate. In some embodiments, thesubstrate is formed in the shape of a helmet within which the head of aperson can be placed, and the grooves are formed as a plurality ofcircumferential grooves or “rings” around a circumference of the helmetfrom the top to the bottom. In some embodiments, the plurality of ringsare separated using a same spacing from the top to the bottom of thesubstrate to create a plurality of levels within which a conductor maybe wound. The plurality of grooves may also include a plurality ofconnecting grooves connecting the circumferential grooves. In someembodiments, the grooves may be formed in the substrate as part offabricating the substrate (e.g., using an additive manufacturingprocess) such that a separate act of forming grooves in the substrate isnot required.

Process 2100 then proceeds to act 2114, where a first portion of aconductor is wound within a first part of the grooves formed in thesubstrate. As discussed above in connection with FIG. 17A, in someembodiments, a first portion of the conductor may be wound from the topto the bottom of the substrate in grooves located at alternating levelsby skipping every other level. In other embodiments, a first portion ofthe conductor may be wound within part (e.g., half turns) of the grooveson each level while skipping other parts of the grooves on each level.Process 2100 then proceeds to act 2116, where a second portion of theconductor is would within a second part of the grooves formed in thesubstrate. For example, the second portion of the substrate when woundfrom the bottom to the top of the substrate may be wound using the partsof the grooves that were skipped when the first part of the conductorwas wound from the top to the bottom. When winding the second part ofthe conductor from the bottom to the top portions of the second part ofthe conductor may cross over (or under) portions of the first part ofthe conductor wound from the top to the bottom. Any suitable conductorincluding, but not limited to, copper wire and litz wire may be used.The conductor may be single stranded or may include multiple strands ofconductive material. The ends of the conductor may be configured to becoupled to an amplifier to amplify signals recorded by the RF coil whenused in a low-field MRI system for receiving MR signals from an imagedobject.

FIGS. 22A-22L illustrate acts in a process of manufacturing atransmit/receive RF head coil for use in a low-field MRI system inaccordance with some embodiments. FIG. 22A shows that the coil windingstarts from the top of a substrate (e.g., a plastic helmet havinggrooves formed therein) in accordance with the numbered arrows. Forexample, the conductor is arranged (1) in a connecting groove connectingthe top of the substrate to a first circumferential groove. Theconductor is then wound (2) in a clockwise direction around a half turnof the first circumferential groove. FIG. 22B shows that aftercompletion (3) of the half turn of the first circumferential groove, theconductor is arranged (4) in a connecting groove connecting the firstcircumferential groove and a second circumferential groove spacedfurther from the top than the first circumferential groove. Theconductor is then wound (5) around an opposite half turn of the secondcircumferential groove also in the clockwise direction. FIG. 22C showsthe conductor is wound (6) within the second circumferential grooveuntil (7) a connecting groove connecting the second and thirdcircumferential grooves is reached as shown in FIG. 22D. The conductoris then arranged (8) within the connecting groove between the second andthird circumferential grooves. FIG. 22E shows that winding (9) continuesin the third circumferential groove in the clockwise direction around anopposite half turn until (10) a connecting groove connecting the thirdand fourth circumferential grooves is reached as shown in FIG. 22F. Theconductor is then arranged (11) within the connecting groove between thethird and fourth circumferential grooves. FIG. 22G shows that windingcontinues in the half turn pattern described above until the bottommostcircumferential groove is reached. In some embodiments, there is nocrossing of the conductor on a posterior side of the helmet as shown inFIG. 22H.

FIG. 22I shows that after completion of the winding on the bottommostcircumferential groove, winding of the conductor continues from bottomto top within the parts of the circumferential grooves that were skippedin the winding from top to bottom. For example, the conductor is wound(12) in the bottommost circumferential groove, the conductor is arranged(13) and crosses over part of the conductor within the connecting grooveconnecting the bottommost circumferential groove and the circumferentialgroove above it. Winding (14) continues in that circumferential groovein the portion that was skipped in the top to bottom winding. As shownin FIG. 22J, the winding (15) continues until the next connecting grooveis encountered, the conductor is arranged (16) and crosses over theconductor within that connecting groove, then continues (17) in the nexthighest circumferential groove. FIG. 22K shows that the windingcontinues in the same pattern to the top of the substrate after whichthe conductor is cut to finish the coil winding for the transmit/receiveRF coil with interlaced winding, as shown in FIG. 22L. Although thewinding has been described as being in the clockwise direction, itshould be appreciated that the winding may alternatively proceed in thecounter-clockwise direction. Additionally, although the processdescribed in FIGS. 22A-L show winding half turns in each of theplurality of circumferential grooves, it should be appreciated that awinding pattern where every other circumferential groove is skipped fromtop to bottom and is then filled from bottom to top may also be used.

FIG. 23A shows a process for manufacturing a receive-only RF coil usingan interlaced winding pattern in accordance with some embodiments. Asshown, the coil winding starts by arranging (1) the conductor from thetop of the substrate (e.g., a plastic helmet having grooves formedtherein) in a connecting groove to one side (e.g., the left side) of thesubstrate. FIG. 23B shows that when the conductor reaches (2) a ringgroove on the left side, the conductor is wound (3) around the ringgroove. FIG. 23C shows that when winding (4) of the conductor around thering groove is completed, the conductor is arranged (5) and crosses overthe conductor in the connecting groove connecting the ring groove andthe top of the substrate. As shown in FIG. 23D, winding (6) continues ina curved groove formed in the left half of the helmet. FIG. 23E showsthat after winding in the left half of the helmet is completed, theconductor is arranged to cross over the top of the helmet to beginwinding of the conductor on the right half of the helmet as shown inFIG. 23F. FIGS. 23G and 23H show that the winding in the right half ofthe helmet continues around the grooves in the right half of the helmetand is arranged and crosses over the conductor in the connecting grooveconnecting the ring groove on the right half of the helmet and the topof the helmet.

Having thus described several aspects and embodiments of the technologyset forth in the disclosure, it is to be appreciated that variousalterations, modifications, and improvements will readily occur to thoseskilled in the art. Such alterations, modifications, and improvementsare intended to be within the spirit and scope of the technologydescribed herein. For example, those of ordinary skill in the art willreadily envision a variety of other means and/or structures forperforming the function and/or obtaining the results and/or one or moreof the advantages described herein, and each of such variations and/ormodifications is deemed to be within the scope of the embodimentsdescribed herein. Those skilled in the art will recognize, or be able toascertain using no more than routine experimentation, many equivalentsto the specific embodiments described herein. It is, therefore, to beunderstood that the foregoing embodiments are presented by way ofexample only and that, within the scope of the appended claims andequivalents thereto, inventive embodiments may be practiced otherwisethan as specifically described. In addition, any combination of two ormore features, systems, articles, materials, kits, and/or methodsdescribed herein, if such features, systems, articles, materials, kits,and/or methods are not mutually inconsistent, is included within thescope of the present disclosure.

The above-described embodiments can be implemented in any of numerousways. One or more aspects and embodiments of the present disclosureinvolving the performance of processes or methods may utilize programinstructions executable by a device (e.g., a computer, a processor, orother device) to perform, or control performance of, the processes ormethods. In this respect, various inventive concepts may be embodied asa computer readable storage medium (or multiple computer readablestorage media) (e.g., a computer memory, one or more floppy discs,compact discs, optical discs, magnetic tapes, flash memories, circuitconfigurations in Field Programmable Gate Arrays or other semiconductordevices, or other tangible computer storage medium) encoded with one ormore programs that, when executed on one or more computers or otherprocessors, perform methods that implement one or more of the variousembodiments described above. The computer readable medium or media canbe transportable, such that the program or programs stored thereon canbe loaded onto one or more different computers or other processors toimplement various ones of the aspects described above. In someembodiments, computer readable media may be non-transitory media.

The terms “program” or “software” are used herein in a generic sense torefer to any type of computer code or set of computer-executableinstructions that can be employed to program a computer or otherprocessor to implement various aspects as described above. Additionally,it should be appreciated that according to one aspect, one or morecomputer programs that when executed perform methods of the presentdisclosure need not reside on a single computer or processor, but may bedistributed in a modular fashion among a number of different computersor processors to implement various aspects of the present disclosure.

Computer-executable instructions may be in many forms, such as programmodules, executed by one or more computers or other devices. Generally,program modules include routines, programs, objects, components, datastructures, etc. that perform particular tasks or implement particularabstract data types. Typically the functionality of the program modulesmay be combined or distributed as desired in various embodiments.

Also, data structures may be stored in computer-readable media in anysuitable form. For simplicity of illustration, data structures may beshown to have fields that are related through location in the datastructure. Such relationships may likewise be achieved by assigningstorage for the fields with locations in a computer-readable medium thatconvey relationship between the fields. However, any suitable mechanismmay be used to establish a relationship between information in fields ofa data structure, including through the use of pointers, tags or othermechanisms that establish relationship between data elements.

The above-described embodiments of the present invention can beimplemented in any of numerous ways. For example, the embodiments may beimplemented using hardware, software or a combination thereof. Whenimplemented in software, the software code can be executed on anysuitable processor or collection of processors, whether provided in asingle computer or distributed among multiple computers. It should beappreciated that any component or collection of components that performthe functions described above can be generically considered as acontroller that controls the above-discussed function. A controller canbe implemented in numerous ways, such as with dedicated hardware, orwith general purpose hardware (e.g., one or more processor) that isprogrammed using microcode or software to perform the functions recitedabove, and may be implemented in a combination of ways when thecontroller corresponds to multiple components of a system.

Further, it should be appreciated that a computer may be embodied in anyof a number of forms, such as a rack-mounted computer, a desktopcomputer, a laptop computer, or a tablet computer, as non-limitingexamples. Additionally, a computer may be embedded in a device notgenerally regarded as a computer but with suitable processingcapabilities, including a Personal Digital Assistant (PDA), a smartphoneor any other suitable portable or fixed electronic device.

Also, a computer may have one or more input and output devices. Thesedevices can be used, among other things, to present a user interface.Examples of output devices that can be used to provide a user interfaceinclude printers or display screens for visual presentation of outputand speakers or other sound generating devices for audible presentationof output. Examples of input devices that can be used for a userinterface include keyboards, and pointing devices, such as mice, touchpads, and digitizing tablets. As another example, a computer may receiveinput information through speech recognition or in other audibleformats.

Such computers may be interconnected by one or more networks in anysuitable form, including a local area network or a wide area network,such as an enterprise network, and intelligent network (IN) or theInternet. Such networks may be based on any suitable technology and mayoperate according to any suitable protocol and may include wirelessnetworks, wired networks or fiber optic networks.

Also, as described, some aspects may be embodied as one or more methods.The acts performed as part of the method may be ordered in any suitableway. Accordingly, embodiments may be constructed in which acts areperformed in an order different than illustrated, which may includeperforming some acts simultaneously, even though shown as sequentialacts in illustrative embodiments.

All definitions, as defined and used herein, should be understood tocontrol over dictionary definitions, definitions in documentsincorporated by reference, and/or ordinary meanings of the definedterms.

The indefinite articles “a” and “an,” as used herein in thespecification and in the claims, unless clearly indicated to thecontrary, should be understood to mean “at least one.”

The phrase “and/or,” as used herein in the specification and in theclaims, should be understood to mean “either or both” of the elements soconjoined, i.e., elements that are conjunctively present in some casesand disjunctively present in other cases. Multiple elements listed with“and/or” should be construed in the same fashion, i.e., “one or more” ofthe elements so conjoined. Other elements may optionally be presentother than the elements specifically identified by the “and/or” clause,whether related or unrelated to those elements specifically identified.Thus, as a non-limiting example, a reference to “A and/or B”, when usedin conjunction with open-ended language such as “comprising” can refer,in one embodiment, to A only (optionally including elements other thanB); in another embodiment, to B only (optionally including elementsother than A); in yet another embodiment, to both A and B (optionallyincluding other elements); etc.

As used herein in the specification and in the claims, the phrase “atleast one,” in reference to a list of one or more elements, should beunderstood to mean at least one element selected from any one or more ofthe elements in the list of elements, but not necessarily including atleast one of each and every element specifically listed within the listof elements and not excluding any combinations of elements in the listof elements. This definition also allows that elements may optionally bepresent other than the elements specifically identified within the listof elements to which the phrase “at least one” refers, whether relatedor unrelated to those elements specifically identified. Thus, as anon-limiting example, “at least one of A and B” (or, equivalently, “atleast one of A or B,” or, equivalently “at least one of A and/or B”) canrefer, in one embodiment, to at least one, optionally including morethan one, A, with no B present (and optionally including elements otherthan B); in another embodiment, to at least one, optionally includingmore than one, B, with no A present (and optionally including elementsother than A); in yet another embodiment, to at least one, optionallyincluding more than one, A, and at least one, optionally including morethan one, B (and optionally including other elements); etc.

Also, the phraseology and terminology used herein is for the purpose ofdescription and should not be regarded as limiting. The use of“including,” “comprising,” or “having,” “containing,” “involving,” andvariations thereof herein, is meant to encompass the items listedthereafter and equivalents thereof as well as additional items.

In the claims, as well as in the specification above, all transitionalphrases such as “comprising,” “including,” “carrying,” “having,”“containing,” “involving,” “holding,” “composed of,” and the like are tobe understood to be open-ended, i.e., to mean including but not limitedto. Only the transitional phrases “consisting of” and “consistingessentially of” shall be closed or semi-closed transitional phrases,respectively.

What is claimed is:
 1. A circuit configured to tune a radio frequency(RF) coil coupled to an amplifier of a low-field magnetic resonanceimaging system, the circuit comprising: tuning circuitry coupled acrossinputs of the amplifier; and active feedback circuitry coupled betweenan output of the amplifier and an input of the amplifier.
 2. The circuitof claim 1, wherein the active feedback circuitry includes at least onefeedback capacitor.
 3. The circuit of claim 1, wherein the tuningcircuitry comprises a tuning capacitor coupled across the inputs of theamplifier.
 4. The circuit of claim 1, wherein the tuning circuitrycomprises a tuning/matching network including at least one capacitor andat least one inductor.
 5. The circuit of claim 1, wherein the activefeedback circuitry is configured to provide a first feedback signal 90or 270 degrees out of phase with a center frequency of the RF coil. 6.The circuit of claim 5, wherein a gain of the amplifier is tuned to be90 or 270 degrees out of phase with the center frequency of the RF coil.7. The circuit of claim 5, wherein one or more components of the activefeedback circuitry is configured to produce the first feedback signalbeing 90 or 270 degrees out of phase with the center frequency of the RFcoil.
 8. The circuit of claim 5, wherein the active feedback circuitryis configured to provide a second feedback signal 180 degrees out ofphase with the center frequency of the RF coil.
 9. The circuit of claim1, wherein the active feedback circuitry is configured to provide afeedback signal 180 degrees out of phase with a center frequency of theRF coil.
 10. The circuit of claim 1, wherein the active feedbackcircuitry includes mutual inductive circuitry including at least oneinductor.
 11. The circuit of claim 10, wherein the at least one inductoris configured to be coupled mutually to the RF coil.
 12. A circuitconfigured to tune a radio frequency (RF) coil coupled to an amplifierof a low-field magnetic resonance imaging system, the circuitcomprising: active feedback circuitry coupled between an output of theamplifier and an input of the amplifier to reduce a quality factor ofthe RF coil.
 13. The circuit of claim 12, wherein the active feedbackcircuitry includes at least one feedback capacitor.
 14. The circuit ofclaim 12, wherein the active feedback circuitry is configured to providea first feedback signal 90 or 270 degrees out of phase with a centerfrequency of the RF coil.
 15. The circuit of claim 14, wherein a gain ofthe amplifier is tuned to be 90 or 270 degrees out of phase with thecenter frequency of the RF coil.
 16. The circuit of claim 14, whereinone or more components of the active feedback circuitry is configured toproduce the first feedback signal being 90 or 270 degrees out of phasewith the center frequency of the RF coil.
 17. The circuit of claim 16,wherein the active feedback circuitry is configured to provide a secondfeedback signal 180 degrees out of phase with the center frequency ofthe RF coil.
 18. The circuit of claim 12, wherein the active feedbackcircuitry is configured to provide a feedback signal 180 degrees out ofphase with a center frequency of the RF coil.
 19. A method of tuning aradio frequency (RF) coil coupled to an amplifier of a low-fieldmagnetic resonance imaging system, the method comprising: arrangingtuning circuitry across first and second inputs of the amplifier; andcoupling active feedback circuitry between an output of the amplifierand an input of the amplifier.
 20. The method of claim 19, furthercomprising providing, using the active feedback circuitry, a feedbacksignal 180 degrees out of phase with a center frequency of the RF coil.